Syndecan‐4 and stromal cell‐derived factor‐1 alpha functionalized endovascular scaffold facilitates adhesion, spreading and differentiation of endothelial colony forming cells and functions under flow and shear stress conditions

Abstract Acellular vascular scaffolds with capture molecules have shown great promise in recruiting circulating endothelial colony forming cells (ECFCs) to promote in vivo endothelialization. A microenvironment conducive to cell spreading and differentiation following initial cell capture are key to the eventual formation of a functional endothelium. In this study, syndecan‐4 and stromal cell‐derived factor‐1 alpha were used to functionalize an elastomeric biomaterial composed of poly(glycerol sebacate), Silk Fibroin and Type I Collagen, termed PFC, to enhance ECFC‐material interaction. Functionalized PFC (fPFC) showed significantly greater ECFCs capture capability under physiological flow. Individual cell spreading area on fPFC (1474 ± 63 μm2) was significantly greater than on PFC (1187 ± 54 μm2) as early as 2 h, indicating enhanced cell–material interaction. Moreover, fPFC significantly upregulated the expression of endothelial cell specific markers such as platelet endothelial cell adhesion molecule (24‐fold) and Von Willebrand Factor (11‐fold) compared with tissue culture plastic after 7 days, demonstrating differentiation of ECFCs into endothelial cells. fPFC fabricated as small diameter conduits and tested using a pulsatile blood flow bioreactor were stable and maintained function. The findings suggest that the new surface functionalization strategy proposed here results in an endovascular material with enhanced endothelialization.

showed significantly greater ECFCs capture capability under physiological flow. Individual cell spreading area on fPFC (1474 ± 63 μm 2 ) was significantly greater than on PFC (1187 ± 54 μm 2 ) as early as 2 h, indicating enhanced cell-material interaction. Moreover, fPFC significantly upregulated the expression of endothelial cell specific markers such as platelet endothelial cell adhesion molecule (24-fold) and Von Willebrand Factor (11-fold) compared with tissue culture plastic after 7 days, demonstrating differentiation of ECFCs into endothelial cells. fPFC fabricated as small diameter conduits and tested using a pulsatile blood flow bioreactor were stable and maintained function. The findings suggest that the new surface functionalization strategy proposed here results in an endovascular material with enhanced endothelialization.

| INTRODUCTION
Myocardial infarction attributed to coronary artery disease (CAD) remains the leading cause of death worldwide. 1 Percutaneous coronary intervention and coronary artery bypass graft (CABG) surgeries are the most common treatments for CAD patients and remains the best practice for patients with complex multivessel disease. 2 While saphenous vein and internal mammary artery grafts are gold standard for CABG surgery, problems associated with these autologous grafts include invasive harvesting procedure and limited availability.
Expanded polytetrafluoroethylene (ePTFE, Teflon ® ) and polyethylene terephthalate (Dacron ® ) grafts have been successfully used clinically in large diameter arterial procedures, but have poor performance in small diameter (<6 mm) CABG surgeries due to the risk for compliance mismatch, thrombosis and neointimal hyperplasia. 3 Consequently, the unmet clinical need for compatible and reliable vascular grafts for CABG surgeries drives the development of novel endovascular biomaterials.
The rapid formation of an intact functional endothelium is critical for the clinical success of any cell-free graft once implanted, especially for small diameter grafts. Endothelial colony forming cells (ECFCs) are of prime importance in vascular remodeling and repair and are directly involved in the formation a functional endothelium. 4 Recent efforts and strategies have been directed toward functionalizing biomaterial surfaces with capture molecules to enhance ECFCs binding capacity since ECFCs circulate at small numbers in the blood. 5 Antibody-based strategies (including antibodies against CD34, CD31, vascular endothelial growth factor receptor 2 [VEGFR2] or vWF) have been shown to facilitate the binding of ECFCs to scaffold materials. 6,7 Although this approach has been used clinically to accelerate endothelization, 8,9 controversy regarding safety and efficacy remains. Intimal hyperplasia has been observed at the venous anastomosis of anti-CD34-coated ePTFE grafts after 4 weeks in a porcine arteriovenous model. 10 In another study, increased density of anti-VEGFR2 antibody functionalized surfaces resulted in an inhibition of cell proliferation. 11 Similarly, although ECFC-specific peptides, aptamers or oligosaccharides have been used as capture molecules, they are not active in promoting cell proliferation and differentiation.
Alternatively, growth factors can be immobilized onto the biomaterial in a physiological manner to recruit ECFCs. It is well known that growth factors are involved in the natural process to mobilize ECFC during vascular injury and promote in situ vascular tissue repair and regeneration. 12 In addition, the immobilization of growth factors on biomaterials has been reported to facilitate the binding of ECFCs as well as promote cell proliferation and differentiation. 13 One growth factor Stromal cell-derived factor-1 alpha (SDF-1α) plays a major role in the homing of circulating progenitor cells through the interaction with CXC chemokine receptor 4 (CXCR4). SDF-1α has been reported to regulate integrin mediated ECFC adhesion to extracellular matrix and differentiation into endothelial cells. 14 Several groups have reported that the functionalization of biomaterial scaffolds with SDF-1α successfully recruited ECFCs and promoted vascular tissue repair. [15][16][17] Clinically available vascular grafts (Gelsoft™ and Polymaille ® C) functionalized with SDF-1α have been shown in ovine models to attract ECFCs, improve endothelialization and reduce intimal hyperplasia and thrombosis. 18 However, the capture rate of ECFCs was maximum at early phase (after 24 h), and less than 50% of the scaffold was covered with endothelium after 3 months. This was possibly due to the adsorbed SDF-1α being nonresilient to pulsatile vascular flow or subjected to enzymatic degradation. Another explanation could be that multiple signaling molecules are involved in promoting cell proliferation and differentiation. Local delivery of a single growth factor may appear to be beneficial initially but may not provide the optimum solution in the long term. To achieve sufficient endothelization on scaffold material, a multifunctional and sustainable approach for functionalizing scaffold material may be beneficial.
In this study, syndecan-4 having multiple ligand binding sites was covalently linked on a biomaterial in order to deliver and attract growth factors to promote endogenous regeneration. This approach is different from existing strategies using either direct adsorption or chemically linked growth factors and may create a unique solution and meet the need for in situ supply of multiple signaling molecules.
The molecular diversity of the oligosaccharide sequences on syndecan-4 makes it a unique candidate to functionalize a variety of scaffold materials and deliver tissue-specific biochemical cues to repair and regenerate different types of tissues. 19 For instance, the binding sites that are specific for SDF-1α, bone morphogen proteins, fibroblast growth factors and vascular endothelial growth factors (VEGF) have been identified in syndecan-4. 20,21 We hypothesized that the combination of syndecan-4 and SDF-1α may work as a functional layer on biomaterial scaffolds to facilitate the binding of ECFCs (see Figure 1). Moreover, syndecan-4 has binding sites to retain and present cell-secreted growth and signaling molecules that are essential for cell proliferation and differentiation. Thus, the syndecan-4 and SDF-1α functionalization of the biomaterial may cause enhanced cell spreading and accelerate cell proliferation and differentiation.
In a previous study, an endovascular biomaterial composite of poly(glycerol sebacate), Silk Fibroin and Type I Collagen (PFC) was created. The composition and nanoscale structure of PFC is similar to extracellular matrix of native vascular tissue and has been shown to support the growth of mature endothelial cells eventually forming a confluent monolayer. 22 PFC is elastic and mechanically durable with a slow degradation rate, which permits vascular regeneration and remodeling in vivo without weakening mechanical properties. 22 In the presented study, PFC was used as a model scaffold material to determine if functionalization with syndecan-4 and SDF-1α would enhance the interaction with ECFCs. It was hypothesized that the functionalization would promote rapid endothelialization by creating a local microenvironment conducive for ECFC adhesion, growth, and differentiation.

| MATERIALS AND METHODS
Poly(glycerol-sebacate) (PGS) prepolymer was synthesized with sebacic acid purchased from Sigma-Aldrich (St. Louis, MO) and glycerol purchased from Fisher Scientific (Waltham, MA) using a published protocol. 23 Type I collagen was purchased commercially from Elastin Products Company, Inc (Owensville, MO). Silk fibroin was extracted from raw silk purchased from Haian Silk Company (Nantong, China) according to published methods. 24 Human bone marrow derived CD34+ cells were purchased from ATCC (Manassas, VA).

| Scaffold fabrication and functionalization
PFC was composed of silk fibroin, type I collagen and PGS at a mass ratio of 4.5:4.5:1. All components were dissolved in HFIP at 10% w/v ratio for electrospinning as published previously. 22 PFC polymer solution was stirred overnight and then loaded into a ing at 120 rpm for 60 min. After electrospinning, the material was removed from the collector and incubated in the oven at 120 C for 48 h to polymerize PGS. Then the material was treated with 1.5% glutaraldehyde vapor overnight to crosslink the proteins. PFC was then functionalized with syndecan-4 and SDF-1α and termed fPFC. 25 Briefly, PFC was saturated with 0.8 μg syndecan-4/cm 2 PFC using the two-step NHS/EDC method and rinsed with Na 2 HPO 4 and PBS. 25 The scaffolds were washed three times with PBS and saturated with 0.8 μg SDF-1α/cm 2 PFC for 2 h at 37 C with agitation. The presence of syndecan-4 and SDF-1α on PFC was confirmed using ELISA.

| ECFC spreading and morphology
Cell spreading of ECFCs on fPFC was compared with PFC and tissue culture plastic (TCP) (N = 3). PFC and fPFC electrospun fiber covered glass coverslips (15 mm in diameter) were fitted into a F I G U R E 1 Schematic illustrating the potential initial interaction of syndecan-4/SDF-1α with circulating ECFCs (Top) and later syndecan-4 interaction with molecules (paracrine and/or autocrine) over time. The figure illustrates early events in endothelialization. Not drawn to scale, only a few syndecan-4 molecules are illustrated.
24-well ultra-low attachment plate. 50,000 ECFCs were added to the plates and cultured in EBM-2 medium supplemented with 2% fetal bovine serum (FBS). After incubation at 37 C for 24 h or 4 days, cells were fixed with 4% PFA, and stained with Rhodamine Phalloidin and SYTOX green. The morphology of ECFCs on the materials was evaluated using a Keyence fluorescence microscope (Itasca, IL). Cell spreading area was quantified using the BZ-X800 Viewer software (version 1.1.2.4). Cell spreading at 2 h was quantified using multiple phase contrast images taken from different fields (N = 9). Violin plots were generated to determine the frequency of cell sizes within a population of cells. For cells cultured for 4 days, the averaged cell spreading area (ACSA) was quantified using multiple fluorescence images (N = 9) and calculated using the following equation: ACSA ¼ Total cell spreading area Total number of cells .

| Cell coverage study
To compare cell coverage rates, 10,000 ECFCs were seeded onto PFC or fPFC electrospun mats (6 mm in diameter) and cultured for 7 days at 37 C (N = 3). Each sample was then removed from the wells, washed with PBS, and fixed with 2.5% glutaraldehyde. Then samples were dehydrated using a series of ethanol solution (50%, 60%, 75%, 80%, 90%, 95%, 100%) and further dried using hexamethyldisilazane from Thermo Scientific (Waltham, MA), followed by coating with a thin layer of gold. Cell morphology and coverage (N = 3) on each sample were evaluated using scanning electron microscopy with the same imaging parameters (Zeiss GeminiSEM 300) and was analyzed with ImageJ.

| ECFC capture study
The capture of ECFCs by scaffolds was evaluated under static and flow conditions. For static conditions, the dynamic binding method was used. 17 PFC and fPFC electrospun mats (6 mm in diameter) were cut and fitted to a 96-well ultra-low attachment plate.
τ w , wall shear stress; μ, viscosity of the medium; Q, flow rate; a, channel height; b, channel width.
ECFCs were seeded at a density of 100,000 cells/cm 2 under either 1 or 10 dynes/cm 2 and interacted with fPFC for 4 h in the flow system. 29 The surface was rinsed to remove unbound cells. The number of ECFCs bound to PFC and fPFC was quantified and normalized using PicoGreen™ DNA quantification.

| Binding specificity of SDF-1α to CXCR4
The binding specificity of ECFCs to fPFC was assessed using the CXCR4 antagonist AMD3100 to compete with SDF-1α (N = 3). For this, ECFCs were incubated with AMD3100 (10 μg/ml) for 30 min to block SDF-1α binding to the CXCR4 receptors. 30

| Selective binding of ECFCs by fPFC
To study SDF-1α binding specificity for ECFCs, a mixed population of peripheral blood leukocytes was used to evaluate the binding of ECFCs to fPFC. Leukocytes were isolated from fresh pig blood using a Results for cells on fPFC and PFC were normalized to gene expression levels on TCP.

| fPFC conduit under physiological arterial flow condition
For these studies fPFC was fabricated as vascular conduits with approximately 4 mm inner diameter (ID; Figure 2A). The conduits were hydrated with PBS for 30 min before study and cut into sections of 3 cm in length and inserted into a cannula at both ends (N = 3).
The ends of the conduit were secured to the cannulas using black silk sutures (Ethicon) to ensure tight connections (see Figure 2B). The ex vivo bioreactor system was kept in an incubator at 37 C and consisted of a flow reservoir (top shelf as shown in Figure 2C tem. 32 The ID of fPFC vascular conduits was measured using a Chison ECO5 Portable Ultrasound immediately after the flow was started and at 1, 3, 6, and 24 h. After the flow cycle was complete, the vascular conduits were retrieved. The dynamic radial compliance (C) was calculated from the ultrasound ID measurements using the following equation:

| Statistical analysis
Group means adopted from preexperiment or estimated from published studies and a power of 0.8 were used to calculate sample sizes.
Statistical analyses were performed using Prism software (version 9.2.0) with either a Student's t-test or a one-way analysis of variance depending on study design. Results were presented as mean ± standard error of the mean for each experiment group. If the results were significant (p < .05), a Tukey's post hoc test was conducted to separate differences in means.

| RESULTS
Initial studies were designed to evaluate the interaction of ECFC with electrospun fibers to determine if the addition of syndecan-4 and SDF-1α to PFC influenced cell attachment and spreading. For these studies, ECFCs were cultured on a sparse layer of fibers to easily monitor the morphology. As shown in Figure 3A cells has been reported to be associated with cell differentiation as well as mechanical transduction. 33 To quantitatively evaluate the influence of fPFC in the initial interaction and spreading of ECFCs, measurements of cell areas were completed using additional randomly selected sets of images like those shown by Figure 3. Violin plots were used to provide the density distribution of ECFC area after 2 h of cell-material interaction ( Figure 4). The wider regions of the plot indicate a higher probability, and the thinner regions represented a lower probability for cell areas.
The median and third quartile of cell spreading areas were greater for fPFC compared with PFC, which suggested a higher population of ECFCs were distributed at the regions of greater spreading areas.
The average cell spreading area of ECFCs cultured on fPFC was 24% greater (p < .01) than for cells on PFC. fPFC had an average cell spreading area of 1474 ± 63 μm 2 and PFC was 1187 ± 54 μm 2 (mean ± SEM).
To further evaluate cell spreading beyond 2 h, the cell morphology and spreading were quantified from multiple sets of fluorescent images after 4 days of culture. There was a significant difference with PFC. The total fraction of area covered by the cells was 47% ± 1.9% for PFC and 59% ± 2.5% for fPFC (p < .01). These observations agreed with previous observations on nanofiber coated coverslips demonstrating that cell surface area coverage was greater for cells cultured on fPFC (see Figure 3). Collectively, the findings suggest  To determine if the SDF-1 bound on PFC was resilient to vascular flow and pressure, conduits of the appropriate size of vessels that typically used in CABG surgeries were fabricated using fPFC. These conduits were tested in a custom-built bioreactor. 32 The initial ID of fPFC conduit was 4 mm. After being conditioned in the pulsatile flow bioreactor for 24 h, the ID did not change with alternating 80 and 120 mmHg pressure as shown in Figure 9A. With over 80,000 relaxation and stress cycles, no deformation, or signs of weakened mechanical properties such as wall bulging or weakening were observed.
Elastic modulus of fPFC was 1.68 ± 0.08 MPa and 1.70 ± 0.03 MPa before and after being conditioned for 24 h ( Figure 9B). Dynamic radial compliance of commercially available Dacron and PTFE grafts has been reported to be 1.9 ± 0.3 and 1.6 ± 0.2 (%/100 mmHg). 37 fPFC had a compliance of 13.7 ± 0.4 (%/100 mmHg), which was closer to internal mammary artery (11.5 ± 3.9 (%/100 mmHg) 38  Previous studies have shown that PFC is compositionally and mechanically similar to native vascular tissue and supports endothelial cell growth in culture. 22 We have also found in a previous study that syndecan-4 and SDF-1α functionalized PFC (fPFC) significantly facilitated the binding of ECFCs compared with syndecan-4 fPFC, SDF-1α fPFC or non-fPFC. 25 In the current study, we confirmed and extended these findings to demonstrate more extensive cell spreading and differentiation of ECFCs if PFC was functionalized with syndecan-4 and SDF-1α. While only SDF-1α was the exogenous growth factor used, the presence of syndecan-4 on PFC may facilitate growth, spreading, and differentiation of ECFCs in multiple and versatile ways (see Figure 1). The molecular domain and structure of sydecan-4 is similar to that found in endothelial glycocalyx, which coordinates the interaction with numerous cellular growth and signaling molecules and maintains hemostatic balance. 39 As a heparan sulfate containing proteoglycan, syndecan-4 has structural diversity to bind various molecules. Heparan sulfate oligosaccharides with specific binding sequences for over 400 molecules have been identified, and include a variety of growth factors, chemokines, cytokines, and anticoagulation molecules. 20,21,40 The presence of syndecan-4 on fPFC may function as a reservoir to capture and present a variety of cell-secreted signaling and growth promoting molecules to surrounding cells, which will positively influence cell-cell and cell-extracellular matrix interaction. This is important for creating a local microenvironment conducive to endogenous regeneration instead of local delivery of stem cells or cell-loaded scaffolds. These procedures typically suffer from drawbacks associated with high cost, prolonged and laborious ex vivo culture process, potential recipient rejection or permanent use of immunosuppressive therapies. 41,42 In the present study, ECFCs on fPFC were morphologically more For patients with CADs, the average WSS over the plaques has been reported to be 42% higher than the healthy region. 49 For CABG surgery, WSS is greater at the anastomotic site and is relatively low along the bed of the graft (up to 11 dynes/cm 2 ). 50  Mature endothelial cells also contribute to the endothelial lining on acellular scaffold material. Known as transanastomotic ingrowth, the host intima grows from anastomosis sites toward the implanted graft material as a response to the injury generated. 54 Biomaterial surfaces such as those of fPFC favoring cell-material interaction may promote vascular repair and regeneration processes by facilitating adhesion, proliferation, and differentiation of a variety of other cell types associated with endothelialization.
Engineering biomaterials that will function and adapt to high arterial pressure and pulsatile flow are among the major challenges in vascular tissue engineering. 55 In order to simulate conditions under which fPFC might be expected to perform upon eventual clinical use, tubular conduits were fabricated and tested. It is well known that saphenous vein grafts used in CABG surgeries dilate after being exposed to high arterial pressure. The fPFC conduits remained intact with no dilations or aneurysmal formations after being exposed to high pressure and repeated cycles of stress and relaxation in an ex vivo pulsatile flow bioreactor. In addition, the dynamic compliance of fPFC indicated it was similar to native internal mammary arteries.
As opposed to commercialized Dacron and PTFE grafts having compliance mismatches with native vascular tissue. These results suggested that fPFC conduits were mechanically durable and able to withstand vascular pulsatile pressures following clinical use.
Native vascular extracellular matrix is formed from a complex organization of fibrous proteins and proteoglycans. 54 The highly porous and interconnected microfibrillar nanostructure not only provides spatial residency and structural support for endogenous cells, but also serves as a reservoir to retain environmental cues that are essential in guiding cellular recruitment, proliferation, and differentiation. Many acellular vascular biomaterial scaffolds are fabricated using electrospinning fabrication protocols. The electrospun nanofiber 3D structures and alignments are very similar to the nanoscale dimension of native vascular extracellular matrix, which make electrospun biomaterials mechanically durable and resistant to vascular flow and pressure. 56 In addition, the stability of the functionalized layer provides the strength of binding and thus the ability to withstand physiologically environment. The results of the current study have shown that ionically bound SDF-1α remains on the material and is resilient to vascular pressure and flow. This finding was in agreement with the high binding affinity of SDF-1α to heparan sulfate (K d = 30 nM) reported in literature. 57 While this study used ECFCs to evaluate the functionalized biomaterial for endovascular repair, these findings can be translated to other scaffolds used thus relate to regenerate a variety of tissues.

DATA AVAILABILITY STATEMENT
The data that supports the findings of this study are available in the supplementary material of this article.